Reflectometry and ellipsometry are light analysis techniques that measure a change in reflective index or polarization of light reflected from a surface of a sample and analyze the measured value to determine a thickness or optical property of the sample.
A reflectometer and an ellipsometer are meters using these techniques. They are used to assess a thickness of various thin films at nano-scale and properties thereof in a nano-thin film manufacturing process of semiconductor industry. Furthermore, the use of them has been expanded to bio-industries and efforts are made to apply to interfacial analysis of biomaterials such as proteins, DNAs, viruses, new drug materials, and the like.
A conventional reflectometer is sufficient to assess a thickness and property of a nano-thin film having a size more than few nanometers (nm), but it has a problem of low reliability due to low detection sensitivity in analysis of a low molecular weight biomaterial which needs sensitivity in the range of about 10.001 nanometers. An ellipsometer has detection sensitivity not more than 0.01 nm compared to the reflectometer, and particularly it has high detection sensitivity under a condition with a large reflective index difference such as measuring a thickness of an oxide film having relatively smaller reflective index compared to a semiconductor onto a semiconductor substrate having a high reflective index.
However, a measurement method having enhanced sensitivity is needed to analyze a low molecular weight biomaterial using the ellipsometer.
As a conventional technique for enhancing detection sensitivity in analysis of biomaterials, there is a surface plasmon resonance sensor (hereinafter, it is referred to as ‘SPR sensor’) in which the reflectometry is mixed with surface plasmon resonance (SPR).
The surface plasmon resonance (SPR) is that electrons present on a metal surface are excited by light waves to cause collective vibration in a normal direction of the surface and absorb light energy. It has been known that the SPR sensor can measure a thickness of a nano-thin film adjacent to the metal surface and a change in reflective index using surface plasmon resonance phenomenon sensitive to polarization of light, as well as allow real-time measurement on a change in adsorption concentration of biomaterials in a non-labeling manner without a fluorescent material.
The SPR sensor is made of a material such as glass coated with a metal thin film in few tens of nanometers, onto which the sensor is attached; the sensor may be bonded with a biomaterial. When a sample in a buffer solution is bonded with the sensor, a resonance angle is changed. The resonance angle is determined by measuring a reflective index. When light enters the SPR sensor, the glass material serves as an incident medium. The light passes through the thin film bonded with the biomaterial, and ultimately the buffer solution serves as a substrate.
In such a configuration, a bio-thin film layer shows a change upon binding with a measured sample, and a shift in resonance angle is directly affected by a refractive index of a buffer solution acting as a substrate. Therefore, to measure pure junction kinetics, the refractive index of the buffer solution should be independently measured and corrected.
To correct a change in reflective index of the buffer solution and avoid an error due to diffusion between the buffer solution and the sample, a method such as using a delicate valve device, an air injection device, and two or more channels and using one channel as a reference channel for correction has been proposed. However, it is difficult to distinguish a SPR angle shift by the reflective index change of the buffer solution from a SPR angle shift by pure adsorption and dissociation property, which may act as a factor causing an error in measurement. Consequently, the conventional SPR sensor shows substantially a difficulty in measuring adsorption and dissociation property of low molecular weight materials due to the foregoing limitation.
Furthermore, the conventional SPR sensor employs a metal thin film made of a noble metal such as gold (Au), silver (Ag) and the like, thereby increasing production costs. Such a metal thin film has uneven roughness according to manufacturing processes and exhibits a large deviation in refractive index. Also, it is difficult to quantitatively measure biomaterials due to unstable optical property. In addition, the SPR sensor comprises an error due to different sensitivities at different positions in relatively comparing with a reference channel.
To improve the foregoing disadvantages of the SPR sensor, a biomaterial binding sensor layer is formed onto the substrate such as silicon, and an amplitude and a phase of light reflected from the substrate through the buffer solution is measured at the P-wave antireflection condition under an immersion microchannel environment using ellipsometry. In this way, a signal of which the measured amplitude is not sensitive to the refractive index change of the buffer solution, but is sensitive to the junction kinetics of the biomaterial can be obtained. Contrary to the SPR measurement, when the junction property of the biomaterial adsorbed on the substrate is measured under an immersion microchannel environment, the buffer solution serves as an incident medium, and light passing through the biomaterial adsorption layer is reflected from the substrate.
In this measurement condition, a measured ellipsometric angle is not sensitive to the refractive index change of the buffer solution serving as an incident medium, but is sensitive to a change of the bio-thin film and the substrate. When a substrate material such as silicon having a stable refractive index is used, since a signal that a measured ellipsometric angle is only sensitive to a change of the bio-thin film can be obtained, the foregoing problems of the SPR method can be solved. However, this method may be applied only when the refractive index change of the buffer solution is negligibly smaller than an ellipsometric angle shift by binding with the bio-thin film. If the refractive index change of the buffer solution is significantly larger than the ellipsometric angle shift by binding with the bio-thin film, the refractive index change should be measured and corrected.
When a solvent having a high refractive index is used for dissolving a sample in a buffer solution or an additional solution is added in the buffer solution to improve surface junction characteristics, one needs a new method for simultaneously and independently measuring the refractive index of the buffer solution and the biomaterial junction characteristics to correct the refractive index of the buffer solution
FIG. 1 shows a configuration of the prior patent wherein a sensor layer is formed onto a substrate such as silicon and measurement is performed under an immersion microchannel environment using ellipsometry to improve the foregoing disadvantages of the SPR sensor. As shown in FIG. 1, the sensor for measuring biomaterial junction characteristics according to the prior patent comprises roughly a microchannel structure unit (100), a substrate (120), a cover part (140), a microchannel (150), a sample injection part (200), a polarized light-generating part (300) and a polarized light-detecting part (400). The sensor for measuring biomaterial junction characteristics according to the prior patent has an adsorption layer (160) formed onto the substrate (120) or a dielectric thin film (130), and forms an immersion environment of the microchannel (150). When a buffer solution (210) containing a biomaterial sample (1) dissolved therein is injected into the microchannel (150), the biomaterial is adsorbed on a ligand material (2) formed on a surface of the adsorption layer (160) to form an adsorption layer having a desired thickness.
Then, polarized incident light from the polarized light-generating part (300) enters a boundary surface between the buffer solution (210) and the substrate (120) through an incident surface (142) at an angle of a P-wave anti-reflection condition. Light reflected from the substrate (120) contains optical data on the adsorption layer of the sample (1). That is, the molecular adsorption and dissociation kinetics such as an adsorption concentration, a thickness of the adsorption layer or a refractive index are changed during the adsorption and dissociation process of the sample (1) to the ligand (2), and hence a measured ellipsometric angle are shifted. The reflected light containing optical data is detected in the polarized light-detecting part (400) through the buffer solution (210) and a reflective surface (144). The polarized light-detecting part (400) measures a change in polarized components of the reflected light, thereby determining the molecular adsorption and dissociation kinetics of the sample (1).
FIG. 2 shows an adsorption curve illustrating an adsorption process of a sample (32) to a metal thin film (20) and a dissociation curve illustrating a dissociation process. A larger adsorption rate constant (ka) means faster absorption of a biomaterial, and a smaller dissociation rate constant (kd) means slower dissociation.
In other words, a dissociation constant (KD=kd/ka) in equilibrium can be calculated by determining the adsorption rate constant and the dissociation rate constant. For example, it can be determined whether a low molecular weight material such as a new drug candidate for a cancer inhibitor may be practically used as a new drug by measuring adsorption or dissociation property of the material to a protein containing a cancer inducer.
Hereinafter, the features and the limitations of the sensor for analyzing biomaterials according to the prior patent will be described with reference to FIG. 3 and FIG. 4. FIG. 3 shows a graph of ellipsometric constants Ψ and Δ on light entering the immersion microchannel structure at the perpendicular in the sensor according to the prior patent. As in FIG. 1, a silicon substrate and a light source (40) having 655 nm wavelength were used.
4 nm and 5 nm thicknesses of a bio-thin film adsorption layer including a thickness of a self-assembled monolayer were respectively measured. The adsorption layer having the refractive index (n) of 1.45 was measured, and buffer solutions having the refractive index (n) of 1.333 and 1.334 were measured. As shown in FIG. 3, a change of Ψ due to the thin film thickness change is relatively larger than a change of Ψ due to the refractive index change. Thus, it can be seen that most Ψ values are caused by the thickness variation.
This largely solves problems caused by a large refractive index change due to a buffer solution in the conventional SPR method, but does not simultaneously measure and correct the refractive index of the buffer solution. This method has an advantage that pure junction kinetics can be measured when a biomaterial sample on which the junction characteristics are measured do not exhibit a large change in refractive index in dissolving it in the buffer solution.
However, if there is a sample which is not well dissolved in the buffer solution, a solvent having a large refractive index difference such as dimethyl sulfoxide (DMSO) may be used, or other materials having different refractive indexes may be added in the buffer solution to increase an electrostatic binding ability and control a pH value. Where a large change in refractive index exhibits by mixing other solvents or materials having different refractive indexes with the pure buffer solution, the refractive index of the buffer solution should be simultaneously measured and corrected. In case that the buffer solution exhibits such large refractive index change, since a change due to the thin film thickness and the refractive index of the buffer solution are together contained in a measured signal, a signal shift by the buffer solution is not negligible.
It can be seen from FIG. 3 that the change of Ψ by perpendicular incident light is directly sensitive to the biomaterial junction kinetics, but the change of Δ is little changed by the thickness or the refractive index change of the buffer solution at an angle which meets the P-wave anti-reflection condition.
FIG. 4 is a schematic view illustrating a problem of the conventional technique that inherent adsorption kinetics of a sample during an adsorption and dissociation process of the sample is mixed with the refractive index change of a buffer solution. FIG. 4(a) is a graph of an inherent adsorption and dissociation concentration of the sample (32). FIG. 4(b) is a graph illustrating a change in measured result depending on the refractive index change of the buffer solution (34). FIG. 4 (c) is a graph of an adsorption and dissociation concentration of the sample (32) measured from the sensor when the inherent adsorption kinetics of the sample (32) is mixed with the refractive index change of the buffer solution. That is, a signal of the sensor is sensitive to an effect of the refractive index change of the buffer solution (34) (arrow), but the adsorption and dissociation concentration of only the sample (32) is unclear. Accordingly, there is a problem that it is difficult to analyze this result and calculate the adsorption and dissociation concentration of the sample (32).
This problem is caused when the refractive index change of the buffer solution during injecting the sample is relatively larger than the signal shift in the sensor by the biomaterial junction characteristics. In particular, this problem is caused when a solvent having a large refractive index difference from the buffer solution is used in mixing the sample with the buffer solution, or other materials used for increasing binding effectiveness has a large refractive index difference from the buffer solution. The conventional SPR sensor shows the foregoing problem even when the refractive index change is minor during injecting the sample. Further, even if substantially the same solutions are injected, the SPR signal shift is significantly larger than the biomaterial junction characteristics, acting as a fundamental factor causing a measurement error.
To correct the refractive index change of the buffer solution (34) and avoid an error due to diffusion between the sample (32) and the buffer solution (34), a method such as using a delicate valve device, an air injection device, and two or more channels and using one channel as a reference channel for correction has been proposed. However, it is different to distinguish the refractive index change of the buffer solution from a change by pure adsorption and dissociation property, and it may always act as a factor causing a measurement error. Consequently, due to said limitations of measurement using the conventional sensor, when the refractive index change of the buffer solution is relatively larger during injecting the sample, there is a fundamental difficulty in measuring dissociation property.
Furthermore, the conventional method can correct a signal varying with time and space only when the reference channel and other channels have constant sensor sensitivity, and a minor change in sensitivity may act as a measurement error factor. In particular, in the SPR sensor, a metal thin film such as gold (Au), silver (Ag) and the like has high production costs, a large deviation in refractive index according to manufacturing processes compared to a semiconductor substrate such as silicon, and unstable optical property. Accordingly, the SPR sensor has a problem that an error is caused due to different sensitivities at different positions in relatively comparing with a reference channel.